As described more detailed herein below, thermal energy deposition is increasingly used in medicine as a means of necrosing diseased tissues. The present invention is disclosed in the following in the context of therapeutic thermal treatment by high intensity focused ultrasound (HIFU). In HIFU a phased array of piezoelectric transducers is used for generating a focused ultrasound beam. It has to be noted, however, that the technique of the invention can be applied equally well in connection with any type of device for the targeted deposition of thermal energy. Examples are lasers or radio frequency antennas.
A therapeutic system comprising an ultrasound therapy unit and a MR imaging unit is generally known, e.g., from WO 2008/152542 A2. In the known system, the MR imaging unit is used to monitor hyperthermia treatments by the ultrasound therapy unit.
Ultrasound is becoming more and more a desirable approach for specific therapeutic interventions. In particular, the use of high intensity focused ultrasound is currently being used as an approach for thermal therapeutic intervention for uterine fibroids and has been examined for possible uses in the treatment of liver, brain, prostate, and other cancerous lesions. Ultrasound has also been the subject of much research as a means for mediating clot dissolution (sonothrombolysis), and has been shown to increase the efficacy of existing medical treatments such as the use of tPA for stroke patients. Ultrasound mediated drug delivery and gene therapy is a further active area of research. Genetic expression of proteins in gene therapy, and increased delivery of drugs in site-targeted therapies have potential to treat a wide variety of diseases with minimal side-effects. Another application for ultrasound therapy is non-invasive treatment for cosmetic means, e.g., removal of fat. The use of ultrasound in all of these applications is desirable because it allows the non-invasive treatment of deep tissues with little or no effect on overlying organs.
In ultrasound therapy for tissue ablation a tissue of interest is insonified with high intensity ultrasound that is absorbed and converted into heat, raising the temperature of the tissue. As the temperature rises above 55° degrees centigrade, coagulative necrosis of the tissue occurs resulting in immediate cell death. The transducers used in therapy can be outside the body or be inserted into the body e.g. through blood vessels, urethra, rectum etc.
MR thermometry, based on the proton resonance frequency shift (PRFS) in water, is presently considered the ‘gold standard’ for the non-invasive monitoring of ablative thermal therapies. The temperature dependence of the proton resonance frequency is primarily due to temperature-induced rupture, stretching, or bending of the hydrogen bonds in water. The temperature dependence of pure water is 0.0107 ppm per degree centigrade, and the temperature dependence of water-based tissues is close to this value. Because of a non-homogeneous magnetic field within the MR imaging apparatus used, absolute proton resonance frequency measurements are not possible. Instead, changes in the proton resonance frequency are measured by first taking a MR image before the delivery of heat, and subtracting this base line thermographic image from subsequent measurements. The temperature-induced changes in the proton resonance frequency are estimated by measuring changes in phase of the MR signal, or frequency shift, by means of appropriate and per se known MR imaging sequences.
Problems arise in applications in which the ultrasound transducer is moved to apply therapy at different locations. The motion of the transducer induces variations in the local magnetic field. The phase images before and after the movement cannot be subtracted to calculate the temperature values. One way to avoid this problem is to wait a sufficiently long time after each motion of the ultrasound transducer in order to allow the tissue to cool down to the baseline value (e.g. 37 degrees centigrade) before further treatment. Then a new baseline thermographic MR image can be collected before sonication begins at the new position and/or orientation of the ultrasound transducer. The drawback of this scheme is that the duration of the treatment is much longer than would actually be necessary.
The presently used MR thermometry sequences do not allow the acquisition of volumetric temperature information in three-dimensional space and for different instances in time. Instead, MR thermometry is presently limited to two-dimensional image planes, thereby enabling reasonable temporal update periods for monitoring the treatment. The location of the image planes for MR thermometry must be carefully chosen. This is because safety must be ensured so that critical anatomic structures and normal tissues are protected. Moreover, it has to be made sure that the intended region has been sufficiently heated and the tissue is completely ablated. In applications in which the ultrasound transducer needs to be moved and treatment has to continue with no interruptions between the sonications, as it is the case, e.g. in intracavitary applications, in which rotational movements of the transducer occur, the image planes for MR thermometry need to be continuously moved an updated. Since the therapy involves a plurality of ultrasound transducer positions and orientations, the image planes used for temperature monitoring cannot be chosen to be present corresponding to all relevant positions and orientations of the transducer.